Inherently de-coupled sandwiched solenoidal array coil

ABSTRACT

An inherently de-coupled sandwiched solenoidal array coil (SSAC) is disclosed for use in receiving nuclear magnetic resonance (NMR) radio frequency (RF) signals in both horizontal and vertical-field magnetic resonance imaging (MRI) systems. In its most basic configuration, the SSAC comprises two coaxial RF receive coils. The first coil of the array has two solenoidal (or loop) sections that are separated from one another along a common axis. The two sections are electrically connected in series but the conductors in each section are wound in opposite directions so that a current through the coil sets up a magnetic field of opposite polarity in each section. The second coil of the SSAC is disposed (“sandwiched”) between the two separated solenoidal sections of the first coil in a region where the combined opposing magnetic fields cancel to become a null. Due to the winding arrangement and geometrical symmetry, the receive coils of the array become electromagnetically “de-coupled” from one another while still maintaining their sensitivity toward receiving NMR signals. The multiple coil array arrangement also allows for selecting between a larger or smaller field-of-view (FOV) to avoid image fold-over problems without time penalty in image data acquisition. Alternative embodiments are disclosed which include unequal constituent coil diameters, unequal constituent coil windings, non-coaxial coil configurations, a three-coil quadrature detection (QD) SSAC arrangement, multiple SSAC arrangements, and optimized SSAC configurations for breast imaging in both horizontal and vertical-field MRI systems.

FIELD OF THE INVENTION

[0001] This invention relates generally to methods and apparatus formagnetic resonance imaging. In particular, it relates to anelectromagnetically de-coupled sandwiched solenoidal array coil forreceiving radio frequency magnetic resonance signals in an MRIapparatus.

BACKGROUND AND SUMMARY OF THE INVENTION

[0002] Based on principals of nuclear magnetic resonance (NMR), magneticresonance imaging (MRI) has become a widely accepted medical imagingmodality—having evolved tremendously over the last two decades as animportant clinical technique for obtaining visual images of tissues andorgan structures within the human body. Basically, clinical MRI relieson the detection of NMR signals from abundant hydrogen protons in thehuman body. These protons are first subjected to a strong radiofrequency (RF) electromagnetic wave excitation pulse. If the frequencyof the excitation pulse is properly chosen, the protons receive neededRF energy to make a transition to an excited state. Eventually, theexcited protons give up their excess energy via a decay process,commonly known as “relaxation”, and return to their original state.

[0003] Since the magnetic moment of a proton is a vector quantity, themicroscopic behavior of millions of protons considered together isequivalent to the vector sum of the individual magnetic moments of allthe protons. For convenience, this sum is typically represented as asingle resultant magnetization vector, M₀, that is aligned with{overscore (B)}₀ (the static main magnetic field). The strong RFexcitation pulse used in MRI effectively tips this resultantmagnetization vector away from alignment with the static main field{overscore (B)}₀ and causes it to precess before decaying back to anequilibrium alignment with {overscore (B)}₀. The component of thisprecessing resultant magnetization vector in a plane perpendicular to{overscore (B)}₀ induces an RF signal, referred to as the nuclearmagnetic resonance (NMR) signal, in an RF “pick-up” or “receive” coil(s)placed near the body portion containing the excited protons.

[0004] During clinical MRI, the magnetic resonance of protons indifferent tissues within an anatomical region are made distinguishablethrough the evocation of a magnetic field gradient along each of threemutually orthogonal spatial directions—the effect of which is to causeprotons at different spatial locations to have slightly differentnuclear magnetic resonance frequencies. The NMR signals induced in thereceive coil can then be processed to reconstruct images of theanatomical structure of interest (i.e., images of the spatialdistribution of NMR nuclei which, in many respects, conform to theanatomical structures containing such nuclei).

[0005] To obtain the maximum induced signal in a receive coil, themagnetic field of the receive coil—conventionally designated as{overscore (B)}₁—must be oriented perpendicular to the direction of thestatic main magnetic field ({overscore (B)}₀) of the MRI apparatus. Fora planar-loop (i.e., a substantially flat loop) type receive coil, thatdirection is in a direction normal to the plane of the conductiveloop(s) of the coil. For a quadrature detection (QD) type coil—whichbasically consists of two RF receive coils having mutuallyperpendicularly oriented {overscore (B)}₁ fields—must also have the{overscore (B)}₁ fields of both of its coils oriented perpendicular tothe MRI apparatus static field {overscore (B)}₀ to obtain a maximuminduced signal.

[0006] Due to the unique nature of the clinical MRI environment, thereare certain design considerations that are particularly relevant towardobtaining maximum performance from an RF receive coil. For example, theNMR signals induced in an RF receive coil during magnetic resonanceimaging are nominally on the order of nanovolts in magnitude while thebackground ambient electrical noise may be of comparable levels orhigher. Therefore, a high performance RF receive coil for clinical MRIneeds to be electromagnetically “sensitive” enough to detect the lowlevel NMR signals despite the relatively high levels of backgroundelectrical noise. Moreover, other design considerations such asfield-of-view, uniformity (i.e., uniformity of the magnetic fieldgenerated by the coil) and coil efficiency are also highly relevant tocoil performance in the clinical MRI environment X coil uniformitybecause it can affect image interpretation and coil efficiency because ahighly efficient coil allows the same image signal information to beacquired within a shorter time frame.

[0007] Theoretical analysis and experimental results have indicated thatfor many MRI applications using multiple RF receive coils together as asignal receiving array is advantageous for improving coil sensitivity,signal-to-noise ratio and imaging field-of-view. Conventionally, theimaging “field-of-view” (FOV) for an MRI receive coil is defined as thedistance indicated between two points on the coil sensitivity profile(i.e., a graph of coil sensitivity vs. distance profile) where thesignal drops to 80% of its peak value. In a typical MRI receive coilarray arrangement, instead of using a single large FOV but lesssensitive coil that covers the entire imaging volume of interest,multiple small FOV but sensitive coils are distributed as an array overthe entire imaging volume. Each individual coil of the array covers asmall localized volume and the NMR signals received by each coil aresimultaneously acquired through corresponding data acquisition channels.Signals from each of the channels are then appropriately combined andprocessed to construct an image of the complete volume of interest. Dueto this ability to simultaneously acquire a signal from multiple sources(i.e., multiple coils) and since each individual signal channel isprovided with its own associated detection circuitry, an array type coilcan conceivably operate with high efficiency. However, the simultaneousacquisition of a signal from a plurality of individual receive coilsnecessitates that each coil function independently, free of interactionor coupling.

[0008] As two individual coils having the same resonance frequency arebrought in close proximity to each other, the common resonance frequencystarts to split into two separate frequencies due to the electromagneticinteraction or “coupling” between the coils. Generally, the closer thecoils are brought together, the stronger the interaction and the largerthe frequency split. Since an MRI receive coil should have its maximumsensitivity optimized for a particular relatively narrow band offrequencies, the resonance frequency splitting can cause sensitivitydegradation when two or more receive coils are closely arranged in anarray.

[0009] Generally, MRI systems are categorized as either a horizontalfield type or vertical field type, based on the direction of the staticmain magnetic field. In a horizontal field system, the static mainmagnetic field is typically oriented in a superior-inferior directionrelative to a patient laying in a prone/supine position. In a verticalfield MRI system, the static main magnetic field is oriented in aninterior-posterior direction relative to a prone/supine patient. Thisdifference in main field orientation is important in that it affects theultimate placement and configuration of an RF receive coil(s) used fordiagnosis in such systems. More often than not, a receive coil designedspecifically for use in a horizontal field system will not be suitablefor similar use in a vertical field system and vice versa.

[0010] Consequently, horizontal field MRI systems and vertical field MRIsystems typically require radically different RF receive coilconfigurations to obtain the maximum achievable performance from thecoil. For example, a planar-loop type receive coil configurationdesigned for obtaining images of a the human spine works well in ahorizontal field MRI system when placed in posterior contact with theback of a patient in supine position. However, the same coilconfiguration may not work in a vertical field system because, in thatcase, the RF magnetic field, {overscore (B)}₁, of the receive coil(i.e., the direction normal to the plane of the loop) is orientedparallel, rather than being perpendicular, to the direction of thestatic main magnetic field {overscore (B)}₀.

[0011] Likewise, a coplanar-loop type array coil is fairly effective inhorizontal-field MRI systems but is rather ineffective in vertical-fieldMRI systems. In a coplanar loop type array coil, the planar-loop receivecoils are arranged in a basically coplanar fashion and distributed overthe imaging volume of interest. Each individual receive coil in acoplanar loop type array is typically a relatively small but highlysensitive RF coil that receives NMR signals from a specific smallportion of the entire region of interest. A final composite image isconstructed by combining signals obtained from each of the individualcoils.

[0012] Magnetic interaction between adjacent coils can be analyzed interms of induced current, induced EMF or magnetic flux. For the purposeof this discussion, a magnetic flux representation is most convenient.In this representation, two coils in close proximity are considered to“couple” to one another if one coil induces a net nonzero magnetic fluxlinkage to the other, and vice versa. Likewise, two coils are consideredto be magnetically “de-coupled” if one coil induces a net zero magneticflux linkage to the other. Consequently, by definition, a completenulling of the magnetic flux linkage between coils in an arrayeffectively “de-couples” the individual coils from one another.

[0013] Since it has become known that magnetic coupling between adjacentelemental coils in a coplanar loop type array coil can be effectivelynulled by properly overlapping the constituent elemental coils (see, forexample, “The NMR Phased Array”, P. B. Roemer et al., Magnetic ResonantMedicine, 1990, 16, pp.192-225), various methods and schemes foroverlapping the elemental coils have contributed toward making coplanararray coils practical and popular for use in horizontal-field MRIsystems. Unfortunately, the coplanar array coils have not been usedsuccessfully in vertical-field MRI systems. Although adapting a coplanararray coil configuration to a vertical-field MRI system has beenattempted by others, such array coils typically have poorsignal—primarily due to the constraint that the normal to the of theconstituent coils (i.e., the direction of the {overscore (B)}₁ field)must be positioned perpendicular to the {overscore (B)}₀ static field ofthe MRI apparatus to obtain the maximum induced signal.

[0014] In an attempt to address this problem, various modifications tothe basic coplanar coil array configuration have been proposed byothers—the more familiar modifications being the so-called “FIG. 8”array coil and its variations. Nonetheless, known prior art attempts toutilize a coplanar array coil or its variants in a vertical-field MRIsystem have resulted in severe limitations in terms of coil sensitivity,imaging depth and uniformity over the desired region of interest.

SUMMARY OF THE INVENTION

[0015] The inventors of the present invention realized that a solenoidaltype coil has many inherent characteristics that make it particularlyadvantageous for use in vertical-field MRI systems. For example,solenoid type coils inherently have high sensitivity and uniformity. Inaddition, the cylindrical shape of the coil fits naturally over variousparts of the body such as the head, neck and other extremities and forclinical MRI applications in a vertical-field system, the B₁ field of asolenoidal type coil when fitted to a patient laying horizontal will beoriented perpendicular to the vertical magnetic field—as required formaximizing the induced signal strength.

[0016] Although it is has been suggested by others that two solenoidalcoils be used in a sandwiched arrangement within an NMR spectrometer toachieve mutual isolation between two coils (see U.S. Pat. No. 4,093,910to Hill, issued Jun. 6, 1978), the two coils comprising this sandwichedarrangement are not used together to form a single RF receiving “array”type antenna for the purpose of obtaining an increase in antennasensitivity or FOV to provide improved image quality for MRI.Specifically, in the NMR spectrometer, one coil is used as a controlchannel RF resonance pick-up coil and the other is used as a sampleanalysis resonance pick-up coil. Moreover, the spectrometer coils areeach separately tuned to be responsive to different resonant frequencies(i.e., one coil is tuned for the sample under analysis and the othercoil is tuned for a control nuclei, e.g., 2D or 19F). In contrast, areceive coil or receive array coil designed for MRI applications shouldhave its maximum sensitivity optimized for a single relatively narrowband of frequencies.

[0017] Accordingly, one general feature of the present invention is aninherently de-coupled array type receive coil having high efficiency,high sensitivity and good uniformity for use in both horizontal andvertical-field MRI systems. The present invention also provides an arraytype receive coil that can be used in both horizontal and vertical-fieldMRI systems to provide an enlarged composite field-of-view (FOV)compared to conventional MRI receive coils. The novel array type receivecoil can be used in horizontal or vertical-field MRI systems to providea capability for allowing selection between different sized FOVs and/orimaging regions. It also provides an array type receive coil that can beused in horizontal or vertical-field MRI systems and which can bereadily adapted to fit one or more disparate size portions of the humananatomy.

[0018] The present invention provides novel RF receive coil arrayarrangements for enhancing the magnetic resonance imaging of one or moreportions of the human anatomy. In particular, the present inventionprovides an inherently magnetically de-coupled array coil for use inmagnetic resonance imaging (MRI) systems. The basic array coil structureof the present invention provides an enhanced signal-to-noise ratio forimproved image quality and a selectable field-of-view (FOV) in bothhorizontal and vertical-field MRI environments. Specifically, thepresent invention utilizes a field-bucking “sandwiched” array coilarrangement that precludes magnetic coupling between constituent coilsof the array while providing increased sensitivity and uniformity overconventional MRI receive coil arrangements.

[0019] In its most basic configuration, the sandwiched array coil of thepresent invention is an array of radio frequency (RF) sensitive coilscomprising at least two solenoidal type RF receive coils arranged suchthat one coil antenna is coaxially surrounded or “sandwiched” betweentwo axially separated solenoidal sections of the second coil, withconductor winding directions being opposite in each of the separatedsections of the second coil. Because of this “sandwiched” structuralconfiguration and the peculiar opposite (field-bucking) windingarrangement, the two RF coils of the array are inherently magnetically“de-coupled” from one another.

[0020] Accordingly, one embodiment of the present invention is a basicsandwiched solenoidal array coil (SSAC) comprising two same-diametercoaxial solenoidal RF coils, each coil having one or more conductivewindings, including an inner coil and an outer coil surrounding theinner coil, the outer coil having separated sections with the conductorwinding directions being opposite in each of the separated sections soas to create a magnetic field bucking arrangement at the position of theinner coil.

[0021] Another embodiment of the present invention is a sandwichedsolenoidal array coil (SSAC) having two non-coaxial, unequal-diameterreceive coils adaptable to fit portions of the human anatomy. Anarrangement for using the SSAC for breast imaging in a vertical-fieldMRI systems for an example non-coaxial, unequal-diameter sandwichedarray coil embodiment is disclosed.

[0022] In a third embodiment of the present invention, a sandwichedsolenoidal array coil (SSAC) is formed wherein the constituentgradient-field coil of the array has solenoidal sections of unequaldiameter and a correspondingly different number of conductive windingsin each section.

[0023] In a fourth embodiment of the present invention, a sandwichedsolenoidal array coil (SSAC) having constituent coils of same ordifferent diameters is adapted for imaging the female breast in ahorizontal-field MRI apparatus.

[0024] In a fifth embodiment of the present invention, a multiplesandwiched solenoidal array coil (SSAC) apparatus is provided forobtaining region-selectable images of the entire human torso.

[0025] In a sixth embodiment of the present invention, a quadraturedetection type sandwiched solenoidal array coil is provided for whichdifferent data channel acquisition arrangements are shown.

BRIEF DESCRIPTION OF THE DRAWINGS

[0026]FIG. 1 is a schematic diagram illustrating the basic configurationof a sandwiched solenoidal array coil in accordance with the presentinvention;

[0027] FIGS. 2A-2E are graphs depicting the {overscore (B)}₁ fielddistributions and the corresponding sensitivity profile for theindividual and combined receive coils of the basic sandwiched solenoidalarray coil of FIG. 1;

[0028] FIGS. 3A-3B are schematic diagrams illustrating the magnetic fluxlinkage between constituent coils of the basic sandwiched solenoidalarray coil of the present invention;

[0029]FIG. 4 is a schematic diagram illustrating a non-coaxial,unequal-diameter sandwiched solenoidal array coil used for breastimaging in a vertical field MRI system in accordance with an alternateembodiment of the present invention;

[0030]FIG. 5 is a schematic diagram illustrating a sandwiched solenoidalarray coil with the gradient-field coil of the array having solenoidalsections of unequal diameters and corresponding different conductivewindings in accordance with an alternate embodiment of the presentinvention;

[0031]FIG. 6 is a schematic diagram illustrating a sandwiched solenoidalarray coil used for breast imaging in a horizontal field MRI apparatusin accordance with an alternate embodiment of the present invention;

[0032]FIG. 7A is a schematic diagram illustrating a region-selectablemultiple SSAC arrangement in accordance with an alternate embodiment ofthe present invention;

[0033]FIG. 7B is a schematic diagram illustrating a single SSAC of themultiple SSAC of FIG. 8A, shown with the connector buckle fastened in aclosed connected position;

[0034]FIG. 8 is a schematic diagram illustrating a quadrature detectiontype sandwiched solenoidal array coil in accordance with an alternateembodiment of the present invention;

[0035] FIGS. 9A-9C are schematic diagrams illustrating three differentdata acquisition arrangements for the quadrature detection SSACembodiment of the present invention.

[0036] FIGS. 10A-10C are example graphs of the electromagnetic isolationobtainable between different coils in the quadrature detection SSACembodiment in accordance with the present invention;

[0037] FIGS. 11A-11C are head phantom images respectively acquired froma quadrature detection coil pair, an RF gradient coil and the basic SSACof the present invention;

[0038] FIGS. 12A-12C are graphs of pixel signal intensity using pixeldata corresponding to the center line, in the superior-inferiordirection, from the image data corresponding to FIGS. 11A, 11B and 11C,respectively; and

[0039]FIG. 13 is a schematic block diagram representation of an exampleMRI system which employs the present invention.

DETAILED DESCRIPTION OF EXEMPLARY EMBODIMENTS

[0040] A diagrammatic representation of the basic sandwiched solenoidalarray coil (SSAC) for receiving NMR signals in accordance with thepresent invention is shown in FIG. 1. In this basic two-coil embodimentof the present invention, receive coil 1 consists of two loop orsolenoidal sections 1 a and 1 b that are spatially separated by adistance W. The two loop or solenoidal sections may consist of eithersingle or multiple conductive windings (turns) and are electricallyconnected by a pair of parallel conductors, 1 c and 1 d, such that anelectrical current (i) in coil 1 flows clockwise in section 1 b butcounterclockwise in section 1 a (or vice versa).

[0041] On its own, the structure of coil 1 is sometimes referred to as a“gradient-field” coil arrangement because the composite magnetic fieldgenerated by the two separated sections changes in magnitude(substantially linearly) as a gradient between the two sections. Inaccordance with the present invention, receive coil 2, which also may bea single-turn loop or a multiple-turn solenoid, is “sandwiched” betweensections 1 a and 1 b of this gradient-field arrangement of coil 1 toform an RF coil array. The separated sections of coil 1 are electricallyconnected and positioned so that a current in coil 1 flows in oppositecircumferential directions through the conductive winding(s) in sections1 a and 1 b such that magnetic fields generated from sections located atopposite sides (axial ends) of the coil 2 achieve a null at the locationof coil 2. This particular structural configuration provides an inherent“decoupling” of the two coils when used together as an array forreceiving NMR signals, as discussed in greater detail below.

[0042] FIGS. 2A-2E show graphs of the {overscore (B)}₁ magnetic fielddistribution and sensitivity profile for each coil of the sandwichedarray, individually, and for the array as a whole. Although B₁ field andcoil sensitivity are two different quantities, they are related. Coilsensitivity represents the ability of a coil in receiving NMR signalfrom a spatial location, whereas B₁ field is the magnetic fieldgenerated at the same spatial location due to unit current flowing inthe coil. The reciprocity principle assures these two effects areproportional (see C-N. Chen et al., “Biomedical Magnetic ResonanceTechnology”, published by Institute of Physics Publishing lop IoppHilger, 1989).

[0043] As depicted in FIG. 2A, the {overscore (B)}₁ field generated bycoil 2 along its longitudinal axial direction, reaches a maximum fieldstrength at the center of coil 2 and drops off as the distance increasesaway from the coil center. Since the magnitude of the {overscore (B)}₁field of a coil is indicative of coil sensitivity, a spatialdistribution of the sensitivity of a coil will have the same generalshape as indicated by a graph of the {overscore (B)}₁ field for thatcoil. As corroborated by the graph of the spatial sensitivity of coil 2shown in FIG. 2B, coil 2 also has its greatest sensitivity very muchlocalized about its center (i.e., where the {overscore (B)}₁ fieldstrength for that coil is maximum).

[0044] As indicated in FIG. 2B, the distance measured between two pointson the coil sensitivity profile where a received NMR signal drops to 80%of its value is, conventionally defined as the 80% field-of-view (FOV)of the coil. Because it is necessary to have a fairly uniformdistribution of coil sensitivity over an area to be imaged to insure auniform signal reception from each point in the imaged area, it isgenerally desirable to have the entire region of interest completelywithin the FOV of the coil.

[0045] Referring now to FIG. 2C, it can be seen that coil 1 has a verydifferent {overscore (B)}₁ field distribution along the central axiscompared to coil 2. At the center of coil 1 and the middle of coil 2,the B₁ fields generated by the two sections, 1 a and 1 b, of coil 1 areequal and opposite—as a result of current flowing in opposite directionin each section—and therefore cancel each other. The {overscore (B)}₁field changes sign (direction) at this middle point and increases instrength toward solenoidal (loop) end sections 1 a and 1 b. (This typeof field distribution is commonly referred to as a “gradient-field” asit changes in magnitude at a substantially linear, constant rate betweentwo points in space.) Since the sign of the B₁ field affects only thephase of the received NMR signal, it is the signal magnitude rather thanthe phase information that is reflected in an MRI magnitude image. Inother words, it is the magnitude of the {overscore (B)}₁ field thatdetermines coil sensitivity. Consequently, the sensitivity of agradient-field coil, such as coil 1, will have an “M” shape profile asevident in FIG. 2D (i.e., two sensitivity “maximums” corresponding tothe two positions of each solenoidal end section, along with a “null”coil sensitivity in the middle).

[0046] When coils 1 and 2 are combined to form a sandwiched array, theresultant signal strength can be weighted average of the individualsignals acquired from each coil. Likewise, the sensitivity profile ofeach coil combines accordingly to produce an overall sensitivity profilefrom which an FOV for the sandwiched array can be derived. In FIG. 2E, agraph of sensitivity versus axial position for coil 1 and coil 2combined together as a sandwiched array is depicted. It can be seen fromthis graph that the sandwiched array coil has an enlarged FOV incomparison with coil 2 alone.

[0047] Referring now to FIGS. 3A and 3B, the magnetic coupling betweenthe coils of the basic SSAC is discussed. For this purpose, it isconvenient to analyze the coils in terms of magnetic flux representativeof the magnetic fields from individual sections. Considering first thecenter “sandwiched” coil 2 of the SSAC, FIG. 3A illustrates the magneticfield, represented by flux lines F, generated by a current flow, i2,flowing in coil 2. Since this magnetic field passes through both loop(or solenoidal) sections 1 a and 1 b of coil 1, the net magnetic fluxlinkage to coil 1 from coil 2 is the sum of the flux linkage to bothsection 1 a and section 1 b. Moreover, since sections 1 a and 1 b aregeometrically symmetrical with respect to coil 2, the flux linkage toeach section is equal in magnitude but opposite in sign. Consequently,the net magnetic flux linkage to coil 1 is zero.

[0048] The above analysis is corroborated by considering the currentinduced in sections 1 a and 1 b of coil 1 by magnetic flux F. Since aninduced current opposes changes in magnetic flux, a current (i1 a)induced in coil section 1 a by magnetic flux F and a current (i1 b)induced in coil section 1 b by the same magnetic flux each flow in thesame relative spacial direction within the individual coil sections.However, due to the opposite conductor winding direction in eachsection, the induced current from each section flows in oppositedirections within the coil as a whole and therefore cancels. As both ofthe preceding methods of analysis confirm, a current flow in coil 2induces no net current flow in coil 1 through magnetic coupling.

[0049] Similarly, referring now to FIG. 3B, if a current 1 is assumed toflow through coil 1, it is evident that current i1 flows clockwise(relatively speaking) in a first coil section 1 b and counterclockwisein a second coil section 1 a. Whereas current i1 flowing in section 1 agenerates a magnetic field indicated by flux Fa, the same current i1flowing in section 1 b generates a magnetic field indicated by flux Fb;and the linkage of magnetic flux from coil 1 to coil 2 will be the sumof the magnetic flux contributed by both coil sections 1 a and 1 b atthe location of coil 2. However, since magnetic flux Fa and Fb are equalin magnitude and opposite in direction at the central “sandwiched”location of coil 2, the two cancel and the total net flux linkage tocoil 2 is zero. Thus, a current flowing in coil 1 induces no net currentin coil 2 via magnetic coupling.

[0050] As evidenced by the preceding discussion with respect to FIGS. 3Aand 3B, an analysis of magnetic flux linkage within the basic SSACbetween constituent coils demonstrates that, due to the geometricalsymmetry and the conductor winding configuration, the two coils areeffectively “de-coupled” from one another. Although, in practice coildimensions and alignment imperfections may result in a slight non-zeronet flux linkage, the magnitude of such residual coupling is small andis easily compensated using conventional decoupling techniques such as asimple capacitor de-coupling network.

[0051] FIGS. 4-7 diagramatically illustrate example structuralconfigurations and some methods of application for four different SSACembodiments in accordance with the present invention. In a firstembodiment, illustrated in FIG. 4, the two RF receiving coils comprisingthe basic SSAC are non-coaxially arranged and each has a differentdiameter. FIG. 4 shows an applied example of this particular embodimentwherein the SSAC of the present invention is configured for imaging ahuman breast of a prone patient in a vertical-field MRI system (asindicated by the directional vectors at the left of the illustration).Despite the non-coaxial configuration and the different diameters, theelectromagnetic analysis discussed above with respect to the basic SSAC(shown in FIGS. 3A and 3B) applies in this case with identical resultsand the two coils can be shown to be electromagnetically “de-coupled.”

[0052] As illustrated in FIG. 4, the SSAC basically consists of a largediameter single-loop or solenoidal coil 2 sandwiched between single-loopor solenoidal sections 1 a and 1 b of a small diameter coil 1. Thelarger diameter coil 2 is placed surrounding both the breast and thetorso while the smaller diameter sections of coil 1 are preferablypositioned “sandwiching” both the breast and coil 2. With thisconfiguration, the SSAC obtains complete coverage of the breast as wellas providing deep penetration for signal sensitivity into the chestwall. In a more accessorized embodiment (not shown here), coil 2 may beconstructed of flexible conductors' and provided with a separableconnector along its top portion, such as a conductive buckle or thelike, for convenient patient loading and adapting to different sizepatients.

[0053] In another embodiment of the SSAC shown in FIG. 5, the basicsandwiched solenoidal array of the present invention is configured suchthat loop (or solenoidal) sections 1 a and 1 b of coil 1 each havedifferent diameters and a corresponding different number of turns (i.e.,conductive windings). Although this particular embodiment of the SSAC iscontemplated primarily for allowing the simultaneous imaging of both thehead and the neck of a patient lying supine in a vertical-field MRIsystem, it can readily be adapted for imaging other portions of theanatomy which may be similarly disparate in size.

[0054] Since the human neck is smaller in diameter than the head, asmaller diameter solenoidal section 1 b is designed for placement aroundthe neck in close proximity to the neck tissues to improve sensitivity.However, since coil section 1 b is smaller in diameter than coil section1 a, more conductive windings (turns) are needed in section 1 b tomaintain a balanced magnetic flux linkage with coil section 1 a. Forexample, as depicted in FIG. 5, a smaller diameter multiple-turndetachable loop section is contemplated for the neck section 1 b of coil1. A larger diameter coil section 1 a and the large diameter“sandwiched” coil 2 are designed to fit over the head of the patient. Inpractice, after a patient is fitted with head section 1 a and coil 2 ofthe SSAC, the detachable neck section 1 b can be fitted over the patientand then reconnected to coil 1, as indicated in FIG. 5 by thecorresponding lettered indicia.

[0055] The basic sandwiched solenoidal array coil of the presentinvention also has practical applications in horizontal-field MRIsystems. For example, an SSAC embodiment contemplated for imaging thehuman breast in a horizontal-field MRI apparatus is illustrated in FIG.6. In this embodiment, a basic SSAC comprising coils 1 and 2 is fittedover and surrounds the breast of a prone patient laying in an MRIapparatus. Sandwiched coil 2 of the SSAC surrounds the center portion ofbreast 3 while sections 1 a and 1 b of coil 1 surround the upper andlower portions. In this manner, the RF magnetic field axis {overscore(B)}₁ of the SSAC is oriented vertically within the horizontal{overscore (B)}₀ field of the MRI apparatus, as indicated in FIG. 6 bythe accompanying vector diagram. Although not shown, coils 1 and 2 ofthe SSAC are connected to appropriate input data acquisition channels ofthe horizontal-field MRI system for combining and processing. Lowersolenoidal section 1 a of coil 1 could also be configured to have asmaller diameter than upper section 1 b so that it can fit closeragainst the breast tissues. In that case, as in the previous embodiment,additional conductive windings (turns) would be needed in the smallerdiameter section to improve signal strength and to balance the magneticflux with section 1 b.

[0056] In accordance with another aspect of the present invention. FIG.7A shows an anatomic-region selectable whole-body coil. In thisembodiment, a multiple SSAC laden “blanket” 6 is designed for wrappingaround a supine patient in a vertical-field MRI system. The blanket ispreferably made up of a soft flexible electrically-insulating materialthat has sets of conductors embedded to form SSAC at multiple positionsalong the length of the blanket when properly connected. The array coilsin the blanket provide NMR signals from distinct major portions of apatient, e.g., head, chest, abdomen, legs and feet.

[0057] Exposed coil conductor portions 4 serve as flexible “belts” forfastening blanket 6 around a patient and each belt is provided with aconductive “buckle” or similar connector 5. Each set of embeddedconductors forming an array coil is also provided with a circuitry andcable (not shown) for delivering NMR signal to data acquisition channelsof the MRI apparatus. In this manner, each set of embedded conductorsforms an array coil that aligns with a different anatomical region of apatient which can then be selected for imaging by fastening the buckleconnectors corresponding to conductive belts of an appropriatelypositioned SSAC, as illustrated in FIG. 7B.

[0058] In an alternate embodiment of the present invention, shown inFIG. 8, a quadrature detection (QD) arrangement is provided byincorporating a saddle coil(s) into the basic SSAC. Saddle coil 3 isco-axially oriented with respect to coil 2 and is centrally locatedalong with coil 2 between the two loop (or solenoidal) sections of coil1. As illustrated by FIG. 8, the direction of the RF magnetic field,{overscore (B)}₁ 3, of saddle coil 3 and the direction of the RFmagnetic field, {overscore (B)}₁ 2, of loop coil 2 are orthogonal toeach other and to the static main field, {overscore (B)}₀, of the MRIapparatus, as indicated by the accompanying magnetic vector diagram.Therefore, coils 2 and 3 form a quadrature detection configuration. Amagnetic flux analysis of this arrangement reveals that coils 1 and 2are electromagntically de-coupled just as in the basic SSAC. Inaddition, due to the geometric symmetry of the arrangement, saddle coil3 is magnetically de-coupled from both coil 1 and coil 2 because its{overscore (B)}₁ field is orthogonal to the {overscore (B)}₁ fieldsthose coils. Any minor residual magnetic coupling between coil 3 andeither of coils 1 or 2 can be compensated using a conventionalcapacitive de-coupling circuit.

[0059] As previously mentioned, the imaging “field-of-view” (FOV) for anMRI receive coil is conventionally defined as the distance between twopoints on the sensitivity profile curve where the signal drops to 80% ofits peak value. Usually, a coil having a large FOV is preferable whenattempting to produce an image that will cover a large region or volumeof interest. However, a large FOV is not particularly desirable whenattempting to obtain an image from just a small region or volume becauseit results in image data “fold-over” problems that can lead to improperimage interpretation. Accordingly, the sandwiched solenoidal array coilof the present invention also permits FOV switching, i.e., switchingbetween a large FOV and a small FOV, without time penalty in image dataacquisition. This means one can reconstruct images with larger FOV orsmaller FOV at the time of data processing, after image data from eachchannel is acquired simultaneously.

[0060] In accordance with another aspect of the present invention, amulti-element (multiple coil) SSAC can be used in either an “array mode”for obtaining simultaneous signal acquisition from each coil in thearray or in a “selectable channel” mode for obtaining a signal from asingle coil or a selected number of coils defining a single dataacquisition channel in the array. For example, using the quadraturedetection embodiment of the present invention shown in FIG. 8,simultaneous signal acquisition from three coils can be accomplishedusing a two-channel arrangement, as shown in FIGS. 9A & 9B. Consideringfirst the arrangement of FIG. 9A, NMR signals picked up by coil 1 areprocessed via a first data acquisition channel 1. At the same time, NMRsignals acquired from coils 2 and 3 are combined using a 90° hybridcombiner and processed via a second data acquisition channel 2. Alimited FOV can be obtained from channel 2 alone or an image having alarger FOV can be obtained by combining signal data from both channel 1and channel 2. A multi-element embodiment of the SSAC thus provides theability to switch between either a large FOV or a small FOV to avoidimage fold-over problems.

[0061] Similarly, other multi-channel signal acquisition and processingarrangements are possible utilizing a multi-element SSAC embodiment ofthe present invention. For example, instead of combining the signalsfrom coils 2 and 3 into a single channel, the signals from coils 1 and 3can be combined, as illustrated in FIG. 9B. Moreover, if a third dataprocessing channel is desirable, the signals from each of the threecoils can be acquired separately in independent channels, as illustratedin FIG. 9C. In addition, each of these signal processing arrangementswill still retain the FOV size switching capability discussed above.

[0062] FIGS. 10A-10C are graphs representing the exemplaryelectromagnetic isolation between different coils in a quadraturedetection SSAC constructed in accordance with the present invention.Each of the three coils in the QD SSAC is connected to its own tune andmatch circuit. Residual electromagnetic coupling between each coil pairis compensated using a variable de-coupling capacitor connected betweeneach coil. The maximum isolation between coils is readily achieved byproperly adjusting the variable de-coupling capacitor. The curves shownin FIGS. 10A-10C represent the S21 parameter using a conventionalnetwork analyzer with the SSAC coils tuned at 15 MHz. FIGS. 10A, 10B and10C illustrate that exemplary electromagnetic isolations between coilsare: −26 dB, −30 dB and −25 dB for coil 1:coil 3, coil 2:coil 3 and coil1:coil 2, respectively.

[0063] FIGS. 11A-11C illustrate example phantom sagital human headimages obtained using the quadrature detection SSAC embodiment of thepresent invention. To obtain these images, an SE 2D imaging sequence wasperformed using a Toshiba Opart vertical imaging system with TR=1000milliseconds, TE=25 milliseconds, FOV=30 cm×30 cm, ST=5 mm and a pixelmatrix of 256×256. FIG. 11A shows a sagital head phantom imageconstructed from signals acquired using coils 2 and 3 as a quadraturedetection pair alone; FIG. 11B shows a sagital head phantom imageconstructed from signals from gradient-field coil 1 alone; and FIG. 11Cshows a sagital head phantom image constructed from the weighted sum ofthe two signals together. The quadrature detection image of FIG. 11Ashows a region of strong signal (i.e., the brightest pixels) at thecenter of the image, whereas the gradient-field coil image of FIG. 11Bshows the strongest signal at the upper and lower edges away from thecenter. As evidenced by FIG. 11C, the image of the weighted sum of thetwo signals together, produced by the SSAC arrangement of the presentinvention, has a much more uniform pixel intensity throughout thanimages from either a quadrature detection coil pair or a gradient-fieldcoil alone.

[0064] FIGS. 12A-12C illustrate the image signal amplitude (pixelintensity) profiles corresponding to data from FIGS. 11A, 11B and 11C,taken from the middle of the respective images, in a superior-inferiordirection. The quadrature detection coil pair alone provides an FOV ofabout 11 centimeters, whereas the quadrature detection SSACconfiguration of the present invention provides an FOV of about 17.5centimeters. Moreover, an even larger FOV may be produced by increasingthe separation, W, between each of the two solenoidal sections of thegradient-field coil of the SSAC.

[0065] Referring now to FIG. 13, a general configuration of the majorcomponents common to a conventional vertical-field MRI system whichemploys the SSAC of the present invention is briefly discussed. Ahorizontal-field MRI system which may also utilize the SSAC of thepresent invention contains the same basic components as thevertical-field system shown, but is structurally configured such thatthe main static polarizing magnetic field, {overscore (B)}₀, is createdsubstantially parallel to the plane of the ground. Typically, an MRIsystem comprises a large superconducting magnet 20 for generating astrong static polarizing magnetic field, an electromagnet coilarrangement 30 for producing large gradient magnetic fields, a gradientmagnetic field pulse generating unit 40, a RF transmitting unit 50 forproducing an RF excitation pulse, a radio frequency receiving unit 60for receiving NMR signals, a signal processing unit 70, a display device80 and a control unit 90 for controlling operation and timing of all theassociated units in the system.

[0066] The static main field generating magnet 20 is arranged in a spacesurrounding patient 10 and provides a powerful uniform vertical (orhorizontal) magnetic field through the patient. An output of thegradient magnetic field generating unit 40 is sent to the coil array 30for producing three gradient magnetic fields G_(x), G_(y) and G_(z)corresponding to X, Y and Z mutually orthogonal directions. An RF pulsegenerated by radio frequency transmitting unit 50 is provided to RF coil51 for transmitting an nuclei excitation pulse into the tissues ofpatient 10. The resulting radio frequency nuclear magnetic resonance(NMR) signals from patient 10 are detected by radio frequency receivingunit 60 after being picked-up by RF receive coil 61. The preferredstructural arrangements for RF receive coil 61 include the abovediscussed example SSAC embodiments of the present invention—which arereadily movable and may be placed on or positioned over specific areasof the patient's body.

[0067] Control unit 90 regulates the timing and application of thegradient magnetic fields and the transmission and reception of RFsignals to RF coils 51 and 61, respectively. An NMR signal output fromradio frequency receiving unit 60 is stored and subjected to Fouriertransform analysis by signal processing unit 70 in a manner well knownin the art to produce image information for displaying on display unit80.

[0068] While the invention has been described in connection with what ispresently considered to be the most practical and preferred embodiment,it is to be understood that the invention is not to be limited to thedisclosed embodiment, but on the contrary, is intended to cover variousmodifications and equivalent arrangements included within the spirit andscope of the appended claims.

What is claimed is:
 1. In an MRI system for producing images from NMRsignals emanating from a subject irradiated with a sequence of RFexcitation pulses in the presence of a polarizing magnetic field andassociated sequences of orthogonal Gradients therein, a sandwichedsolenoidal array coil for receiving NMR signals emanating from a regionof interest in the subject, said array coil comprising: a firstsolenoidal coil having one or more conductive windings; and a secondsolenoidal coil arranged along a common axis with the first coil andhaving at least two electrically connected solenoidal sections, eachsection having one or more conductive windings, disposed along thecommon axis such that at least one solenoidal section is located at eachaxial end of the first coil and a current in the second coil flows inopposite circumferential directions through conductive windings insections located at opposite ends of the first coil so as to create amagnetic field bucking arrangement effective over a region encompassingthe first coil.
 2. The sandwiched solenoidal array coil of claim 1,wherein the magnetic field bucking arrangement of said second coil issuch that magnetic fields emanating from the solenoidal sections becomea null at the center of the first coil.
 3. The sandwiched solenoidalarray coil of claim 1, wherein one or more solenoidal sections of thesecond coil have different diameters and a correspondingly differentquantity of conductive windings such that magnetic fields emanating fromthe solenoidal sections become a null at the center of the first coil.4. The sandwiched solenoidal array coil of claim 1, wherein a diameterof the first coil is different from a diameter of one or more solenoidalsections of the second coil.
 5. In an MRI system for producing imagesfrom NMR signals emanating from a subject irradiated with a sequence ofRF excitation pulses in the presence of a polarizing magnetic field andassociated sequences of orthogonal gradients therein, a sandwichedsolenoidal array coil for receiving the NMR signals emanating from aregion of interest in the subject, said array coil comprising: a firstcoil having one or more conductive windings; a second coil having atleast two spatially separated sections connected by a pair ofconductors, each section having one or more conductive windings,disposed so that at least one section is located at each axial end ofthe first coil such that the first coil is effectively sandwichedbetween two sections of the second coil and the separated sectionselectrically connected such that a current in the second coil flows inopposite circumferential directions through the conductive windings insections located at opposite ends of the first coil so that magneticfields emanating from the sections become a null at the center of thefirst coil.
 6. A sandwiched solenoidal array coil as set forth in claim5, wherein a diameter of the first coil is different from a diameter ofone or more sections of the second coil.
 7. A sandwiched solenoidalarray coil as set forth in claim 5, wherein separated sections of thesecond coil have a different diameter and a correspondingly differentquantity of conductive windings.
 8. A quadrature detection sandwichedsolenoidal array coil arrangement for receiving NMR signals in a MRIapparatus, said array coil comprising: a first solenoidal loop coilhaving one or more conductive windings; at least two further, spatiallyseparated, solenoidal coil sections, each section having one or moreconductive windings, disposed so that at least one section is located atopposite axial ends of the first solenoidal loop coil such that thefirst solenoidal loop coil is effectively sandwiched between the twoseparated sections and magnetic fields emanating from the separatedsections become a null at the center of the first solenoidal loop coil;and a saddle coil coaxially oriented with respect to the firstsolenoidal loop coil and disposed between the solenoidal loop coil, thetwo separated sections, wherein the first solenoidal loop coil, theseparated coil sections and saddle coil together form an array coilwhich is oriented such that the magnetic field axis of the solenoidalloop coil and the magnetic field axis of the saddle coil are orthogonalto each other and to a main static polarizing magnetic field axis, B₀,of the MRI apparatus.
 9. A quadrature detection sandwiched solenoidalarray coil arrangement as set forth in claim 8, wherein the separatedcoil sections provide a first data acquisition channel and the saddlecoil and the solenoidal loop coil are connected via a 90° hybridcombiner to provide a second data acquisition channel.
 10. A quadraturedetection sandwiched solenoidal array coil arrangement as set forth inclaim 8, wherein the separated coil sections and the saddle coil areconnected via a 90° hybrid combiner to provide a first data acquisitionchannel and the solenoidal loop coil provides a second data acquisitionchannel.
 11. A quadrature detection sandwiched solenoidal array coilarrangement as set forth in claim 8, wherein the separated coil sectionsprovide a first data acquisition channel, the solenoidal loop coilprovides a second data acquisition channel, and the saddle coil providesa third data acquisition channel.
 12. A method for imaging a breast of aprone patient laying in an MRI apparatus having a vertical main staticpolarizing magnetic field, said method comprising the steps of: a)fitting a breast of said prone patient into an open axial end of asandwiched solenoidal array coil comprising a plurality of coaxial coilelements wherein at least one coil element is a solenoidal loop coilsandwiched between a separated pair of solenoidal coil sections; b)orienting the sandwiched solenoidal array coil within the MRI apparatussuch that a RF magnetic field. {overscore (B)}₁, produced by thesolenoidal loop coil is substantially orthogonal to the vertical mainstatic polarizing magnetic field of the MRI apparatus; and c) using saidsandwiched solenoidal array coil to receive NMR signals during animaging operation of the MRI apparatus.
 13. A method for imaging abreast a prone patient laying in a horizontal main field MRI apparatus,said method comprising the steps of: a) fitting a patient with asandwiched solenoidal array coil comprising a plurality of coaxial coilelements wherein at least one coil element is a solenoidal loop coilsandwiched between separated solenoidal coil sections and wherein thesolenoidal loop coil of the array is of a large diameter and is fittedover the torso and breast of the patient and the separated coil sectionsof the array are fitted such that the breast is sandwiched between theseparated coil sections; b) orienting the sandwiched solenoidal arraycoil within the MRI apparatus such that a RF magnetic field, {overscore(B)}₁, of the solenoidal loop coil is substantially orthogonal to ahorizontal main static polarizing magnetic field of the MRI apparatus;and c) using said sandwiched solenoidal array coil to receive NMRsignals during an imaging operation of the MRI apparatus.
 14. Aregion-selectable sandwiched solenoidal array coil for imaging the torsoof a patient utilizing an MRI apparatus, comprising: a sheet of flexibleelectrically insulating material having one or more embedded conductorsfor forming a plurality of sandwiched solenoidal array coils, eachsandwiched solenoidal array coil comprising a plurality of array coilelements wherein at least one coil element is a solenoidal loopsandwiched between a pair of solenoidal loop coil sections, wherein suchsandwiched solenoidal array coils are formed surrounding the patient atselected positions along the length of the sheet when the sheet iswrapped around the torso of a patient and conductor portions thereof areselectively connected together.
 15. A region-selectable sandwichedsolenoidal array coil as set forth in claim 14, wherein conductorportions include a conductive belt portion and a conductive buckleportion for electrically and structurally connecting embedded conductorportions together.
 16. A region-selectable sandwiched solenoidal arraycoil as set forth in claim 14, wherein the sheet comprises two or morelaminated layers of material and contains conductors between laminatedlayers.
 17. A region-selectable sandwiched solenoidal array coil as setforth in claim 14, wherein the sheet forms a blanket.
 18. In a MRIsystem having multiple data acquisition channels for producing imagesfrom NMR signals emanating from a subject irradiated with an RFexcitation pulse in the presence of a static polarizing magnetic field,said system having a sandwiched solenoidal array coil for receiving theNMR signals emanating from a region of interest in the subject, saidsandwiched solenoidal array coil comprising a plurality of coil elementswherein at least one coil element is a solenoidal loop coil sandwichedbetween a pair of solenoidal coil sections, a method for selectivelyobtaining differently sized FOVs of an imaged region of interest, saidmethod comprising the steps of: a) providing NMR signals received by aselected first solenoidal loop coil to a first data acquisition channelof the MRI system for processing; b) providing NMR signals received by asandwiching pair of solenoidal coil sections to a second dataacquisition channel of the MRI system for processing; and c) selectingbetween: (i) processing signal data acquired from both the first and thesecond data acquisition channels and combining processed data to producean image or (ii) processing signal data acquired from a selected singledata acquisition channel to produce an image, wherein an image producedfrom combined data acquired from both the first and the second dataacquisition channels results in an FOV of enhanced size compared to animage produced from a single data acquisition channel alone.
 19. Amethod for selectively obtaining differently sized FOVs as set forth inclaim 18 wherein the sandwiched solenoidal array coil includes asaddle-coil element and NMR signals received by the saddle-coil elementare combined with the first coil element.
 20. In an MRI system havingmultiple data acquisition channels for producing images from NMR signalsemanating from a subject irradiated with a sequence of RF excitationpulses in the presence of a static polarizing magnetic field andassociated sequences of orthogonal gradients therein, said system havinga one or more sandwiched solenoidal array coils for receiving the NMRsignals emanating from one or more regions of interest in the subject,said sandwiched solenoidal array coils comprising a plurality of coilelements wherein at least one coil element is a first solenoidal loopcoil and at least one coil element is a split-apart solenoidal loopdisposed to encompass the first solenoidal loop, a method forselectively obtaining images from different regions of interest in thesubject, said method comprising the steps of: a) providing NMR signalsreceived by a first sandwiched solenoidal array coil, selectivelylocated at a first region of interest on a subject, to a first dataacquisition channel of the MRI system for processing; b) providing NMRsignals received by a second sandwiched solenoidal array coil,selectively located at a second region of interest on a subject, to asecond data acquisition channel of the MRI system for processing; and c)selecting between processing signal data acquired from the first dataacquisition channel or the second data acquisition channel to produce animage, wherein an image produced from the first data acquisition channelcovers a different region of interest than an image produced from thesecond data acquisition channel.
 21. In an MRI system having multipledata acquisition channels for producing images from NMR signalsemanating from a subject irradiated with a sequence of RF excitationpulses in the presence of a static polarizing magnetic field andassociated sequences of orthogonal gradients therein, said system havinga sandwiched solenoidal array coil for receiving the NMR signalsemanating from a region of interest in the subject, said sandwichedsolenoidal array coil comprising a plurality of coil elements wherein atleast one coil element is a first solenoidal loop coil and at least onecoil element is a split apart solenoidal coil encompassing the firstelement, a method for selectively obtaining images from differentportions within a region of interest, said method comprising the stepsof: providing NMR signals obtained from a plurality of individual coilelements of the sandwiched solenoidal array coil to a plurality ofindividual data acquisition channels of the MRI apparatus; andselectively switching between processing an NMR signal from a particularcoil element by switching between different data acquisition channels,wherein images from different regions of interest of the subject areobtained from different data acquisition channels.
 22. In a MRI systemhaving multiple data acquisition channels for producing images from NMRsignals emanating from a subject irradiated with a sequence of RFexcitation pulses in the presence of a static polarizing magnetic fieldand associated sequences of orthogonal gradients therein, said systemhaving a sandwiched solenoidal array coil for receiving the NMR signalsemanating from a region of interest in the subject, a method forproducing an image of a region of interest of a subject, said methodcomprising the steps of: simultaneously obtaining an NMR signal from aselected plurality of coil elements and providing said signal via one ormore data acquisition channels of the MRI apparatus for processing toproduce an image of the region of interest.
 23. In a MRI system havingmultiple data acquisition channels for producing images from NMR signalsemanating from a subject irradiated with a sequence of RF excitationpulses in the presence of a static polarizing magnetic field andassociated sequences of orthogonal gradients therein, said system havinga sandwiched solenoidal array coil for receiving the NMR signalsemanating from a region of interest in the subject, a method forsimultaneously imaging the head and neck portions of a human patient,said method comprising the steps of: a) fitting the head and neckportions of a patient with a sandwiched solenoidal array coil comprisingseparated solenoidal elements having different diameters correspondingto respective head and neck portions of a patient, at least someseparated solenoidal sections having a correspondingly differentquantity of conductive windings such that magnetic fields emanating fromthe sections become a null at the center of another at least onesolenoidal loop coil element; b) orienting the sandwiched solenoidalarray coil within the MRI apparatus such that its RF magnetic field,{overscore (B)}₁, is substantially orthogonal to the main polarizingmagnetic field of the MRI apparatus; and c) using said sandwichedsolenoidal array coil to receive NMR signals during an imaging operationof the MRI apparatus.
 24. A method for simultaneously imaging the headand neck portions of a human patient as set forth in claim 23 wherein atleast one of the solenoidal loop coil sections has a diametercorresponding to the head portion of a patient.
 25. A method forsimultaneously imaging the head and neck portions of a human patient asset forth in claim 23 wherein at least one of the solenoidal loop coilsections has a diameter corresponding to the neck portion of a patient.